Apparatus and method for medical image reconstruction using deep learning for computed tomography (ct) image noise and artifacts reduction

ABSTRACT

A method and apparatus is provided that uses a deep learning (DL) network to reduce noise and artifacts in reconstructed medical images, such as images generated using computed tomography, positron emission tomography, and magnetic resonance imaging. The DL network can operate either on pre-reconstruction data or on a reconstructed image. The DL network can be an artificial neural network or a convolutional neural network (e.g., using a three-channel volumetric kernel architecture). Different neural networks can be trained depending on the noise level, scanning protocol, or the anatomic, diagnostic or clinical objective of the reconstructed image (e.g., by partitioning the training data into noise-level range and training respective DL networks for each range). Further, the DL networks can be trained to mitigate artifacts, such as the cone-beam artifact.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.15/727.216, filed on Oct. 6, 2017, the entire contents of which areincorporated herein by reference.

FIELD

This disclosure relates to using deep learning (DL) networks to improvethe image quality of reconstructed medical images, and, moreparticularly, to using DL networks to reduce noise and artifacts inreconstructed computed tomography (CT), positron emission tomography(PET), and magnetic resonance imaging (MRI) images.

BACKGROUND

The background description provided herein is for the purpose ofgenerally presenting the context of the disclosure. Work of thepresently named inventors, to the extent the work is described in thisbackground section, as well as aspects of the description that may nototherwise quality as prior art at the time of tiling, are neitherexpressly nor impliedly admitted as prior art against the presentdisclosure.

Computed tomography (CT) systems and methods are widely used,particularly for medical imaging and diagnosis. CT systems generallycreate images of one or more sectional slices through a subject's body.A radiation source, such as an X-ray source, irradiates the body fromone side. At least one detector on the opposite side of the bodyreceives radiation transmitted through the body. The attenuation of theradiation that has passed through the body is measured by processingelectrical signals received from the detector.

A CT sinogram indicates attenuation through the body as a function ofposition along a detector array and as a function of the projectionangle between the X-ray source and the detector array for variousprojection measurements. In a sinogram, the spatial dimensions refer tothe position along the array of X-ray detectors. The time/angledimension refers to the projection angle of X-rays, which changes as afunction of time during a CT scan. The attenuation resulting from aportion of the imaged object (e.g., a vertebra) will trace out a sinewave around the vertical axis. Those portions farther from the axis ofrotation correspond to sine waves with larger amplitudes, and the phaseof the sine waves corresponds to the angular positions of objects aroundthe rotation axis. Performing an inverse Radon transform—or any otherimage reconstruction method—reconstructs an image from the projectiondata in the sinogram.

X-ray CT has found extensive clinical applications in cancer, heart, andbrain imaging. As CT has been increasingly used for a variety ofapplications including, e.g., cancer screening and pediatric imaging,there has arisen a push to reduce the radiation dose of clinical CTscans to become as low as reasonably achievable, for low-dose CT, theimage quality can be degraded by many factors, such as high quanta noisechallenge scanning geometry (i.e., large cone angle, high helical pitch,truncation, etc.), and other non-ideal physical phenomenon (i.e.,scatter, beam hardening, crosstalk, metal, etc.). Developing efficientcorrection methods can be challenging due to the difficulties ofmodelling accurate forward model and solving complicated inverseproblem.

Although many cutting-edge technologies have been developed during thepast decades to improve low-dose CT image quality, such as model basediterative image reconstruction or sinogram restoration, those methodsare often time consuming and requires expensive hardware. Particularly,at some challenge scenarios, the image qualities are still interior tothe high dose images. Accordingly, improved methods are desired in orderto reduce computational time, hardware costs and further improvelow-dose CT image quality.

BRIEF DESCRIPTION OF THE DRAWINGS

A more complete understanding of this disclosure is provided byreference to the following detailed description when considered inconnection with the accompanying drawings, wherein:

FIG. 1A shows an example of a flow diagram of a method for reducingnoise and/or an artifact that uses a deep-learning (DL) network toprocess a reconstructed image, according to one implementation;

FIG. 1B shows an example of a flow diagram of a method for reducingnoise and/or an artifact that uses a DL network to process sinogramdata, according to one implementation;

FIG. 2A shows an example of a DL network that is a feedforwardartificial neural network (ANN), according to one implementation;

FIG. 2B shows an example of a DL network that is a convolutional neuralnetwork (CNN), according to one implementation;

FIG. 2C shows an example of implementing a convolution layer for oneneuronal node of the convolution layer, according to one implementation;

FIG. 2D shows an example of a implementing a three channel volumetricconvolution layer for volumetric image data, according to oneimplementation;

FIG. 3 shows an example of a How diagram for training a DL network,according to one implementation;

FIG. 4 shows an example of a flow diagram for applying the ANN,according to one implementation;

FIG. 5 shows an example of a flow diagram for applying the CNN,according to one implementation;

FIG. 6 shows a schematic of an implementation of a computed tomography(CT) scanner, according to one implementation;

FIG. 7 shows a schematic of an implementation of a magnetic resonanceimaging (MRI) scanner, according to one implementation;

FIG. 8A shows a perspective view of a positron-emission tomography (PET)scanner, according to one implementation; and

FIG. 8B shows a schematic view of the PET scanner, according to oneimplementation.

DETAILED DESCRIPTION

To address the above-identified challenges of known reconstructionmethods for medical images, the methods described herein have beendeveloped in order to reduce computational time, hardware costs, andfurther improve image quality low-dose medical images, such as computedtomography (CT) images. Further, the examples provided herein ofapplying these methods are non-limiting, and the methods describedherein can benefit other medical imaging modalities such as MRI,PET/SPECT, etc. by adapting the framework proposed herein. Accordingly,the discussion herein discloses and describes merely exemplaryimplementations of the present disclosure. As will be understood bythose skilled in the art, the present disclosure may be embodied inother specific forms without departing from the spirit or essentialcharacteristics thereof. Accordingly, the present disclosure is intendedto be illustrative, but not limiting of the scope of the invention, aswell as other claims. The disclosure, including any readily discerniblevariants of the teachings herein, defines, in part, the scope of theforegoing claim terminology such that no inventive subject matter isdedicated to the public.

In general, it is desirable to reduce CT radiation dose as low asreasonably achievable (ALARA) white maintaining diagnostic quality.Clinical applications for which reduced radiation dose and low-countcomputed tomography (CT) are advantageous include: CT perfusion study,low and ultra-low-dose CT screening, low dose whole body imaging formelanoma or pediatrics, bias noise reduction for lower kVp imaging indual energy CT to reduce total dose, ultra-low-dose CT for PET/CTattenuation correction (CTAC), respiratory-gated CT for phased matchedCTAC, and motion correction for PET.

As discussed above, in low-dose CT, the image quality can be degraded bymany factors, such as high quanta noise challenge scanning geometry andother non-ideal physical phenomenon. As a result, developing efficientcorrection methods can be challenging due to the difficulties ofmodelling an accurate forward model and of solving a complicated inverseproblem. For example, model based iterative image reconstruction orsinogram restoration can be time consuming and require expensivehardware.

To address the above-identified challenges of known methods, the methodsdescribed herein use deep learning (DL) networks. In general, DLnetworks have been adapted to image processing area for improving imagespatial resolution and reducing noise. As compared to the traditionalmethods, deep learning does not require accurate noise and edgemodelling, relying instead on training data sets. Further, deep learninghas the capability to capture the interlayer image features by buildingup a sophisticated network between noisy observations and latent cleanimages.

For example, the methods herein leverage improvements in variousresearch areas whereby DL-based convolutional neural network (CNN) canbe applied to denoising reconstructed images and/or sinogramrestoration. Methods applying DL-based CNN to CT image reconstructionare mostly unknown. Training data corresponding to different CT scanningmethods and scanning conditions can be used to train various CNNnetworks to be tailored for projection data corresponding to particularCT scanning methods, protocols, applications, and conditions by usingtraining data selected to match the particular CT scanning methods,protocols, applications, and conditions. Thus, respective CNN networkscan be customized and tailored to certain conditions and methods for CTscanning. Additionally, the customization of the CNN networks can extendto the noise level or to the signal-to-noise ratio of the projectiondata, and can be extended to the anatomical structure or region of thebody being imaged. The methods described herein can be applied tosinogram restoration in addition to denoising of reconstructed images.Further, the redundancy of information in adjacent slices of athree-dimensional CT image can be used to perform volumetric-based DL byusing a kernel for the convolution layers of the DL network that extendsto pixels in slices above and below a slice that is being denoised. Andin addition to denoising the reconstructed images, the DL can be trainedto mitigate artifacts in the reconstructed images.

In particular, various implementations of the methods described hereinprovide several advantages over previous methods of imagereconstruction. First, certain implementations of the methods describedherein can use DL to optimize the compensation weights andreconstruction filters in FBP algorithm. For example, in certainimplementations, the methods described herein use DL to leverage upanalytical reconstruction and provide images with comparable imagequality as high-dose model based iterative reconstructions.

Second, certain implementations of the methods described herein performoffline training of a DL network and embed the trained network in thereconstruction step. For example, in certain implementations, themethods described herein can exhibit the benefits associated with: (i)using a three-channel based network; (ii) classifying training databased on noise level; and (iii) optimizing the training sets byconsidering/accounting for the anatomical features. Therefore, methodsdescribed herein can have better image quality in terms of lower noiseand higher spatial resolution than previous methods.

Third, the methods described herein can use iterative reconstruction(IR) for offline high-dose training data proration. Accordingly, incertain implementations, clean images similar to those obtained using anIR method can be achieved by applying the DL network to reconstructedimages generated using a lease time-intensive reconstruction method(e.g., filtered back-projection), resulting in a dramatic reduction inthe computational time and expensive hardware while achieving the imagequality of an IR method without the computational burden of an IRmethod.

In general, DL can be adapted to image processing area for improvingimage spatial resolution and reducing noise. As compared to thetraditional methods. DL does not require accurate noise and edgemodelling and only relies on training data sets. Further, DL has thecapability to capture the interlayer image features and build up ansophisticated network between noisy observations and latent cleanimages.

Referring now to the drawings, wherein like reference numerals designateidentical or corresponding parts throughout the several views. FIG. 1Ashows a flow diagram of method 100, which has two process: process 110for offline training and process 140 for reconstructing a high-qualityCT image from projection data (which can also be referred to as asinogram).

The process 110 of method 100 performs offline training of the DLnetwork 135. In step 130 of process 110, noisy data 115 and optimizeddata 120 are used as training data to train a DL network, resulting inthe DL network being output from step 130. More generally, data 115 canbe referred to as defect-exhibiting data, for which the “defect” can beany undesirable characteristic that can be affected through imageprocessing (e.g., noise or an artifact). Similarly, data 120 can bereferred to as defect-reduced data, defect-minimized data, or optimizeddata, for which the “defect” is less than in the data 115. In an exampleusing reconstructed images for data 115 and 120, the offline DL trainingprocess 110 trains the DL network 135 using a large number of noisyreconstructed images 115 that are paired with correspondinghigh-image-quality images 120 to train the DL network 135 to produceimages resembling the high-image-quality images from the noisyreconstructed images.

In process 140 of method 100, the protection data 145 is corrected instep 150, and then, in step 160, a CT image is reconstructed from thecorrected projection data using an image reconstruction process (e.g.,an inverse Radon transformation).

In step 150, the projection data can be corrected for a detector offset(e.g., due to dark current or noise), pile up, variations in quantumefficiency in the detectors (e.g., between detector elements and as afunction of energy of the X-ray photons), etc. Further, thesecorrections can be based on calibration data, empirical, and knownparameters (e.g., the geometry of the scanner, the detector elements,anti-scatter grids, etc.).

In step 160, the image reconstruction can be performed using aback-projection method, a filtered back-projection method, aFourier-transform-based image reconstruction method, an iterative imagereconstruction method (e.g., algebraic reconstruction technique), amatrix inversion image reconstruction method, or a statistical imagereconstruction method.

In step 170, the reconstructed image is denoised using the DL network135. The result of which is a high-quality image 175. Thus, noisy CTimages resulting from the CT reconstruction in step 160 can be processedusing a DL denoising algorithm applying the network generated by theoffline DL training process 110.

FIG. 1B shows an alternative implementation of method 100. In method100′ shown in FIG. 1B, the DL network 135′ is applied in step 170′ torestoring the sinogram before the image reconstruction step 160, ratherthan denoising the reconstructed image after the image reconstructionstep 160. In this ease the DL network 135′ represents a network that hasbeen trained at step 130′ of process 110′ using a large number of noisysinograms 115 that are paired with corresponding high-quality sinograms120. For example, in step 140′, raw data 145 (e.g., pre-log) can beprocessed by pre-log corrections and converted to sinogram data in step150. Then, in the sinogram restoration step 170′ and the reconstructionstep 160, the DL network 135′ is, applied to sinogram restoration, and,after sinogram correction, image reconstructions are applied to generatethe high-quality image 175.

It is also contemplated that in certain implementations a DL network135′ can be used to restore a sinogram and a DL network 135 can be usedto denoise the image reconstructed from the restored sinogram within asingle method 100 to generate the high-quality image 175.

FIGS. 2A, 2B, 2C, and 2D show various examples of the DL network 135(135′).

FIG. 2A shows an example of a general artificial neural network (ANN)having N inputs, K hidden layers, and three outputs. Each layer is madeup of nodes (also called neurons), and each node performs a weighted sumof the inputs and compares the result of the weighted sum to a thresholdto generate an output. ANNs make up a class of functions for which themembers of the class are obtained by varying thresholds, connectionweights, or specifics of the architecture such as the number of nodesand/or their connectivity. The nodes in an ANN can be referred to asneurons (or as neuronal nodes), and the neurons can haveinter-connections between the different layers of the ANN system. Thesimplest ANN has three layers, and is called an autoencoder. The DLnetwork 135 generally has more than three layers of neurons, and has asmany outputs neurons as input neurons, wherein N is the number of pixelsin the reconstructed image (sinogram). The synapses (i.e., theconnections between neurons) store values called “weights” (alsointerchangeably referred to as “coefficients” or “weightingcoefficients”) that manipulate the data in the calculations. The outputsof the ANN depend on three types of parameters: (i) the interconnectionpattern between the different layers of neurons, (ii) the learningprocess for updating the weights of the interconnections, and (iii) theactivation function that converts a neuron's weighted input to itsoutput activation.

Mathematically, a neuron's network function m(x) is defined as acomposition of other functions n_(i)(x), which can further be defined asa composition of other functions. This can be conveniently representedas a network structure, with arrows depicting the dependencies betweenvariables, as shown in FIGS. 2A and 2B. For example, the ANN can use anonlinear weighted sum, wherein m(x)=K(Σ_(i)w_(i)n_(i)(x)), where K(commonly referred to as the activation function) is some predefinedfunction, such as the hyperbolic tangent.

In FIG. 2A (and similarly in FIG. 2B). the neurons (i.e., nodes) aredepicted by circles around a threshold function. For the non-limitingexample shown in FIG. 2A, the inputs are depicted as circles around alinear function, and the arrows indicate directed connections betweenneurons. In certain implementations, the DL network 135 is a feedforwardnetwork as exemplified in FIGS. 2A and 2B (e.g., it can be representedas a directed acyclic graph).

The DL network 135 operates to achieve a specific task, such asdenoising a CT image, by searching within the class of functions F tolearn, using a set of observations, to find m*∈F which solves thespecific task in some optimal sense (e.g., the stopping criteria used instep 260 of step 130 discussed below). For example, in certainimplementations, this can be achieved by defining a cost function C:F→isuch that, for the optimal solution m* ,C(m*)≤C(m)∀m∈F (i.e., nosolution has a cost less than the cost of the optimal solution). Thecost function C is a measure of how far away a particular solution isfrom an optimal solution to the problem to be solved (e.g., the error).Learning algorithms iteratively search through the solution space tofind a function that has the smallest possible cost. In certainimplementations, the cost is minimized over a sample of the data (i.e.,the training data).

FIG. 2B shows a non-limiting example in which the DL network 135 is aconvolutional neural network (CNN). CNNs are type of ANN that hasbeneficial properties for image processing, and, therefore, havespecially relevancy for the applications of image denoising and sinogramrestoration. CNNs use feed-forward ANNs in which the connectivitypattern between neurons can represent convolutions in image processing.For example, CNNs can be used for image-processing optimization by usingmultiple layers of small neuron collections which process portions ofthe input image, called receptive fields. The outputs of thesecollections can then tiled so that they overlap, to obtain a betterrepresentation of the original image. This processing pattern can berepeated over multiple layers having alternating convolution and poolinglayers.

FIG. 2C shows an example of a 4×4 kernel being applied to map valuesfrom an input layer representing a two-dimensional image to a firsthidden layer, which is a convolution layer. The kernel maps respective4×4 pixel regions to corresponding neurons of the first hidden layer.

Following after a convolutional layer, a CNN can include local and/orglobal pooling layers, which combine the outputs of neuron clusters inthe convolution layers. Additionally, in certain implementations, theCNN can also include various combinations of convolutional and fullyconnected layers, with pointwise nonlinearity applied at the end of oralter each layer.

CNNs have several advantages for image processing. To reduce the numberof free parameters and improve generalization, a convolution operationon small regions of input is introduced. One significant advantage ofcertain implementations of CNNs is the use of shared weight inconvolutional layers, which means that the same filter (weights bank) isused as the coefficients for each pixel in the layer; this both reducesmemory footprint and improves performance. Compared to otherimage-processing methods, CNNs advantageously use relatively littlepre-processing. This means that the network is responsible for learningthe filters that in traditional algorithms were hand-engineered. Thelack of dependence on prior knowledge and human effort in designingfeatures is a major advantage for CNNs.

FIG. 2D shows an implementation of DL network 135 that takes advantageof the similarities between adjacent layers in reconstructedthree-dimensional medical images. The signal in adjacent layers isordinarily highly correlated, whereas the noise is not. That is, ingeneral, a three-dimensional volumetric image in CT usually can providemore diagnostic information than single slice transverse two-dimensionalimage since more volumetric features can be captured. Based in thisinsight, certain implementations of the methods described herein use avolumetric-based deep-learning algorithm to improve the CT images. Thisinsight and corresponding method also applies to other medical imagingareas such as MRI, PET, etc.

As shown in FIG. 2D, a slice and the adjacent slices (i.e., the sliceabove and below the central slice) are identified as a three-channelinput for the network. To these three layers, a W×W×3 kernel is appliedM times to generate M values for the convolutional layer, which are thenused for the following network layers/hierarchies (e.g., a poolinglayer). This W×W×3 kernel can also be thought of end implemented asthree W×W kernels respectively applied as three-channel kernels that areapplied to the three slices of volumetric image data, and the result isan output for the central layer, which is used as an input for thefollowing network hierarchies. The value M is the total filter numberfor a given slice of the convolutional layer, and W is the kernel size(e.g., W=4 in FIG. 2C).

In certain implementations, the DL network 135 is not a single networkbut is instead several networks, each optimized for a different set ofconditions of a CT scan. For example, the DL networks 135 can beoptimized according to a series of noise classes, corresponding torespective ranges for the signal-to-noise ratio (SNR) or ranges for thenoise level. The level of noise relative to the signal in a CT imageoften depends on the total penetration photon numbers. That is, ahigh-flux photon number results in higher SNR, whereas a low-flux photonnumber results in a lower SNR.

Accordingly, capturing the image characteristics at different noiselevels can be a beneficial aspect of the offline training process 110,having a significant impact the resulting image quality. Certainimplementations of the methods described herein address the differencesencountered for different SNRs by training different DL networks 135according to different ranges of the noise level, especially forlow-dose CT image reconstruction.

In the offline training process 110, the noisy images are classifiedbased on their noise level ranges. For each class/range, a separatenetwork of the DL networks 135 is trained. That is, the DL networks 135include several networks, each corresponding to a specific noise-levelrange and is trained using noisy images 115 corresponding to thenoise-level range.

Then, in the CT image noise reduction step 170, the noise level of thereconstructed image is measured to determine in which noise-level rangethe reconstructed image belongs. Based on this determination, acorresponding network is selected from the DL networks 135 to performstep 170 for the post processing the reconstructed image. By performingnoise-level based training process, the network can be optimized toeliminate the noise texture and artifacts (for example, the streakartifacts in a low-close scenario) specific to a particular noise level.Ignoring noise level can reduce the image quality and residual undesirednoise features might persist.

Additionally, in certain implementations, different networks can betrained for the DL networks 135 based on the type of anatomic featuresbeing imaged. To better capture and represent the image featurescorresponding to different anatomic structures and/or clinicalapplications, diagnostic images can be optimized using optimized images120 having qualities tailored to respective clinical applications and/oranatomic structures and, in certain implementations, using tailoredoptimization strategies and cost functions. For example, the trainingdata can be categorized based on anatomical structures (e.g., head,abdomen, lung, cardiac, etc.). Further, the training data can be furthercategorized according to anatomical structure. For example, thesespecially tailored DL networks 135 can be trained using speciallygenerated training data by generating, for each anatomical structureand/or clinical or diagnostic application, pairs of noisy andhigh-quality images, which are reconstructed for specific anatomicalstructures and/or with specific reconstruction parameters or kernels forthe given clinical or diagnostic purposes.

Then, in step 130, the categories of training data are used to trainrespective DL networks 135 are trained for different anatomicalstructure and/or clinical applications. Given the trained DL networks135, step 170 performs CT image noise reduction using the appropriatetrained network(s) selected based on anatomy/application.

Although the above variations for implementing method 100 and method100′ have been exemplified using method 100 to denoise reconstructedimages, each variation can also be used with method 100′ to restoresinograms prior to image reconstruction in step 160.

Further, method 100 can be used to mitigate artifacts instead of or inaddition to denoising reconstructed images. For example, large-anglecone-beam CT (CBCT) scans are desirable for applications benefiting fromrapid/short scan times. However, large-angle CBCT can also suffer fromvarious artifacts. When method 100 is tailored for artifact reduction,steps 130 and 170 can be modified as described below.

In step 130, the architecture of the DL network 135 can be modified inorder to be optimized for artifact correction by training the DL network135 using artifact data 115 that exhibits the artifact and artifact freedata 120 that is predominantly free of the artifact. For example, theartifact free data 120 can be generated using a process that is selectedto maintain the artifact below a predetermined threshold. In certainimplementations, the training data includes high-quality images 120 canbe generated using optimized scan condition. (e.g., high dose, smallcone angle, with known internal material, etc.) and the correspondinglow-quality images 115 can be generated using scan conditions exhibitingthe artifact at issue (e.g., the scan conditions and reconstructionprocess that are anticipated/defaults to be used during the scan used togenerate the projection data 145 and the reconstruction method used instep 160). Then the DL networks 135 are trained to remove the artifactsfrom the low-quality images 115 by optimizing the deep learningarchitecture of the DL networks 135.

In step 170, the reconstructed image from step 160 is processed using anappropriate network architecture from the DL networks 135 to output thehigh-quality image 175, which is a high-quality clean image that enablesclinicians to make a diagnosis with greater diagnostic confidence. Theselected DL network 135 has an appropriate network architecture when theconditions and parameters of the scan and reconstruction giving rise tothe reconstructed image generated in step 160 correspond to theconditions and parameters of the data 115 used to train the selected DLnetwork 135.

Consider for example the case of large cone-angle scanning protocols. Toreduce the patients motion artifacts and improve image temporalresolution, large cone-angle scanning protocols are often used for head,cardiac and functional CT images. However, due to the angle limitationof incident x-ray beam, the images at large cone angle position mightnot have sufficient data to reconstruct certain volume pixels within thereconstructed image volume, resulting in artifacts. Although empiricalmethods, such as z-axis interpolation, have been proposed to compensatethe large cone angle problem, these methods fail to completely eliminateartifacts. To better eliminate the artifact at large cone angle, a DLnetwork 135 can be optimized and applied in the image domain to reducethe cone beam artifacts. This can be achieved by selecting the trainingdata 115 and 120 to be pairs of images of the same object or patientscanned, in which each pair includes one image using a large-cone-angleprotocol and one image using a small-cone-angle protocol (e.g., usinghelical scans). That is, the artifact data 115 can be generated usingthe same large-cone-angle protocol that will be used to generate theprojection data 145 (e.g., 320 segments), and the optimized data 120scan be generated using a small-cone-angle protocol (e.g., 80 or fewersegments) that maintains the cone-beam artifacts below a predeterminedthreshold. Then, the DL network 135 will be optimized to correct theartifacts in images reconstructed from projection data 145 generatedusing a large-cone-angle protocol with 320 segments.

FIG. 3 shows one implementation of supervised learning used to train theDL network 135 in step 130. In supervised learning, a set of trainingdata is obtained, and the network is iteratively updated to reduceerror, such that the noisy data 115 processed by the DL network closelymatches the optimized data 120. In other words, DL network infers themapping implied by the training data, and the cost function produces anerror value related to the mismatch between the optimized data 120 andthe denoised data produced by applying a current incarnation of the DLnetwork 135 to the noisy data 115. For example, in certainimplementations, the cost function can use the mean-squared error tominimize the average squared error. In the case of a of multilayerperceptrons (MLP) neural network, the backpropagation algorithm can beused for training the network by minimizing the mean-squared-error-basedcost function using a gradient descent method.

Training a neural network model essentially means selecting one modelfrom the set of allowed models (or, in a Bayesian framework, determininga distribution over the set of allowed models) that minimizes the costcriterion (i.e., the error value calculated using the cost function).Generally, the DL network can be trained using any of numerousalgorithms for training neural network models (e.g., by applyingoptimization theory and statistical estimation).

For example, the optimization method used in training artificial neuralnetworks can use some form of gradient descent, using backpropagation tocompute the actual gradients. This is done by taking the derivative ofthe cost function with respect to the network parameters and thenchanging those parameters in a gradient-related direction. Thebackpropagation training algorithm can be: a steepest descent method(e.g., with variable learning rate, with variable learning rate andmomentum, and resilient backpropagation), a quasi-Newton method (e.g.,Broyden-Fletcher-Goldfarb-Shanno, one step secant, andLevenberg-Marquardt), or a conjugate gradient method (e.g.,Fletcher-Reeves update, Polak-Ribiére update, Powell-Beale restart, andscaled conjugate gradient). Additionally, evolutionary methods, such asgene expression programming, simulated annealing,expectation-maximization, non-parametric methods and particle swarmoptimization, can also be used for training the DL neural networks 135.

FIG. 3 shows a non-limiting example of a flow diagram of animplementation of step 130 of method 100 (and similarly for step 130′ ofmethod 100′) for training the network using the training data. The data115 in the training data can be a noisy image or an image exhibiting anartifact. For example, an artifact can arise from a particular method ofreconstruction, or arise from a method used for acquiring the projectiondata (e.g., a large-angle cone beam acquisition).

In step 210 of step 130, an initial guess is generated for thecoefficients of the DL network 135. For example, the initial guess canbe based on a priori knowledge of the region being imaged or one or moreexemplary denoising methods, edge-detection methods, and/or blobdetection methods. Additionally, the initial guess can be based on a DLnetwork 135 trained on training data related to a different noise levelor using a different CT scan method, as discussed above.

Exemplary denoising methods include linear smoothing filters,anisotropic diffusion, non-local means, or nonlinear filters. Linearsmoothing filters remove noise by convolving the original image with amask that represents a low-pass filter or smoothing operation. Forexample, the Gaussian mask comprises elements determined by a Gaussianfunction. This convolution brings the value at pixel into closeragreement with the values of its neighbors. Anisotropic diffusionremoves noise while preserving sharp edges by evolving an image under asmoothing partial differential equation similar to the heat equation. Amedian filter is an example of a nonlinear filter and, if properlydesigned, a nonlinear filter can also preserve edges and avoid blurring.The median filter is one example of a rank-conditioned rank-selection(RCRS) filter, which can be applied to remove and pepper noise from animage without introducing significant blurring artifacts. Additionally,a filter using a total-variation (TV) minimization regularization termcan be applied if imaged region supports an assumption of uniformityover large areas that are demarked by sharp boundaries between theuniform areas. A TV filter is another example of a nonlinear filter.Moreover, non-local means filtering, is an exemplary method ofdetermining denoised pixels using a weighted average over similarpatches within the images.

In step 220 of step 130, an error (e.g., a cost function) is calculatedbetween the network processed noisy data 115 and the optimized data 120.The error can be calculated using any known cost function or distancemeasure between the image (sinogram) data, including those costfunctions described above.

In step 230 of step 130, a change in the error as a function of thechange in the network can be calculated (e.g., an error gradient), andthis change in the error can be used to select a direction and step sizefor a subsequent change to the weights/coefficients of the DL network135. Calculating the gradient of the error in this manner is consistentwith certain implementations of a gradient descent optimization method.In certain other implementations, as would be understood by one ofordinary skill in the art, this step can be omitted and/or substitutedwith another step in accordance with another optimization algorithm(e.g., a non-gradient descent optimization algorithm like simulatedannealing or a genetic algorithm).

In step 240 of step 130, a new set of coefficients are determined forthe DL network 135. For example, the weights/coefficients can be updatedusing the change calculated in step 230, as in a gradient descentoptimization method or an over-relaxation acceleration method.

In step 250 of step 130, a new error value is calculated using theupdated weights /coefficients of the DL network 135.

In step 260 of step 130, predefined stopping criteria are used todetermine whether the training of the network is complete. For example,the predefined stopping criteria can evaluate whether the new errorand/or the total number of iterations performed exceed predefinedvalues. For example, the stopping criteria can be satisfied if eitherthe new error falls below a predefined threshold or if a maximum numberof iterations is reached. When the stopping criteria is not satisfiedprocess 130 will continue back to the start of the iterative loop byreturning and repeating step 230 using the new weights and coefficients(the iterative loop includes steps 230, 240, 250, and 260). When thestopping criteria are satisfied process 130 is completed.

In addition to the implementation for error minimization shown in FIG. 3, process 130 can use one of many other known minimization methods,including, e.g., local minimization methods, convex optimizationmethods, and global optimization methods.

When the cost function (e.g., the error) has local minima that aredifferent from the global minimum, a robust stochastic optimizationprocess is beneficial to find the global minimum of the cost function.Examples of optimization method for finding a local minimum can be oneof a Nelder-Mead simplex method, a gradient-descent method, a Newton'smethod, a conjugate gradient method, a shooting method, or other knownlocal optimization method. There are also many known methods for findingglobal minima including: genetic algorithms, simulated annealing,exhaustive searches, interval methods, and other conventionaldeterministic, stochastic, heuristic, and metatheuristic methods. Any ofthese methods can be used to optimize the weights and coefficients ofthe DL network. Additionally, neural networks can be optimized using aback-propagation method.

FIGS. 4 and 5 show flow diagrams of implementations of step 170. FIG. 5is general for all ANNs, and FIG. 4 is particular to CNNs. Further,FIGS. 4 and 5 are also applicable to step 170′ with the substitutionthat the DL network operates on sinogram data rather than areconstructed image. The implementation of step 170 shown in FIG. 4corresponds to applying the DL network 135 to an image that has beenreconstructed in step 160.

In step 410, the weights/coefficients corresponding to the connectionsbetween neurons (i.e., nodes) are applied to the respective inputscorresponding to the pixels of the reconstructed image.

In step 420, the weighted inputs are summed. When the only non-zeroweights/coefficients connecting to a given neuron on the next layer areregionally localized in an image represented in the previous layer, thecombination of steps 410 and 420 is essentially identical to performinga convolution operation.

In step 430, respective thresholds are applied to the weighted sums ofthe respective neurons.

In process 440 the steps of weighting, summing, and thresholding arerepeated for each of the subsequent layers.

FIG. 5 shows a flow diagram of another implementation of step 170. Theimplementation of step 170 (170′) shown in FIG. 5 corresponds tooperating on the reconstructed image (sinogram data) using anon-limiting implementation of a CNN for the DL network 135.

In step 450, the calculations for a convolution layer are performed asdiscussed in the foregoing and in accordance with the understanding ofconvolution layers of one of ordinary skill in the art.

In step 460, the outputs from the convolution layer are the inputs intoa pooling layer that is performed according to the foregoing descriptionof pooling layers and in accordance with the understanding of poolinglayers of one of ordinary skill in the art.

In process 470 the steps of a convolution layer followed by a poolingcan be repeated a predefined number of layers. Following (or intermixedwith) the convolution and pooling layers, the output from a poolinglayer can be fed to a predefined number of ANN layers that are performedaccording to the description provided for the ANN layers in FIG. 4 . Thefinal out will be a reconstructed image having the desirednoise/artifact free characteristics.

FIG. 6 illustrates an implementation of the radiography gantry includedin a CT apparatus or scanner. As shown in FIG. 6 , a radiography gantry500 is illustrated from a side view and further includes at X-ray tube501, an annular frame 502, and a multi-row or two-dimensional-array-typeX-ray detector 503. The X-ray 501 and X-ray detector 503 arediametrically mounted across an object OBJ on the annular frame 502,which is rotatably supported around a rotation axis RA. A rotating unit507 rotates the annular frame 502 at a high speed, such as 0.4sec/rotation, while the object OBJ is being moved along the axis RA intoor out of the illustrated page.

The first embodiment of an X-ray computed tomography (CT) apparatusaccording to present inventions will be described below with referenceto the views of the accompanying drawing. Note that X-ray CT apparatusesinclude various types of apparatus , e.g., a rotate/rotate-typeapparatus in which an X-ray tube and X-ray detector rotate togetheraround an object to be examined, and a stationary/rotate-type apparatusin which many detection elements are arrayed in the form of a ring orplane, and only an X-ray tube rotates around an object to be examined.The present inventions can be applied to either type. In this case, therotate/rotate type, which is currently the mainstream, will beexemplified.

The multi-slice X-ray CT apparatus further includes as high voltagegenerator 509 that generates a tube voltage applied to the X-ray tube501 through at slip ring 508 so that the X-ray tube 501 generatesX-rays. The X-rays are emitted towards the object OBJ, whose crosssectional area is represented by a circle. For example, the X-ray tube501 having an average X-ray energy during a first scan that is less thanan average X-ray energy during a second scan. Thus, two or more scanscan be obtained corresponding to different X-ray energies. The X-raydetector 503 is located at an opposite side from the X-ray tube 501across the object OBJ for detecting the emitted X-rays that havetransmitted through the object OBJ. The X-ray detector 503 furtherincludes individual detector elements or units.

The CT apparatus further includes other devices for processing thedetected signals from X-ray detector 503. A data acquisition circuit ora Data Acquisition System (DAS) 504 converts a signal output from theX-ray detector 503 for each channel into a voltage signal, amplifies thesignal, and further converts the signal into a digital signal. The X-raydetector 503 and the DAS 504 are configured to handle a predeterminedtotal number of projections per rotation (TPPR).

The above-described data is sent to a preprocessing device 506, which ishoused in a console outside the radiography gantry 500 through anon-contact data transmitter 505. The preprocessing device 506 performscertain corrections, such as sensitivity correction on the raw data. Amemory 512 stores the resultant data, which is also called projectiondata at a stage immediately before reconstruction processing. The memory512 is connected to a system controller 510 through a data/control bus511, together with a reconstruction device 514, input device 515, anddisplay 516. The system controller 510 controls a current regulator 513that limits the current to a level sufficient for driving the CT system.

The detectors are rotated and/or fixed with respect to the patient amongvarious generations of the CT scanner systems, in one implementation,the above-described CT system can be an example of a combinedthird-generation geometry and fourth-generation geometry system. In thethird-generation system, the X-ray tube 501 and the X-ray detector 503are diametrically mounted on the annular frame 502 and are rotatedaround the object OBJ as the annular frame 502 is rotated about therotation axis RA. In the fourth-generation geometry system, thedetectors are fixedly placed around the patient and an X-ray tuberotates around the patient. In an alternative embodiment, theradiography gantry 500 has multiple detectors arranged on the annularframe 502, which is supported by a C-arm and a stand.

The memory 512 can store the measurement value representative of theirradiance of the X-rays at the X-ray detector unit 503. Further, thememory 512 can store a dedicated program for executing various steps ofmethod 100 and/or method 100′ for correcting low-count data and CT imagereconstruction.

The reconstruction device 514 can execute various steps of method 100and/or method 100′. Further, reconstruction device 514 can executepre-reconstruction processing image processing such as volume renderingprocessing and image difference processing as needed.

The pre-reconstruction processing of the projection data performed bythe preprocessing device 506 can include correcting for detectorcalibrations, detector nonlinearities, and polar effects, for example.Further, the pre-reconstruction processing can include various steps ofmethod 100 and/or method 100′.

Post-reconstruction processing performed by the reconstruction device514 can include filtering and smoothing the image, volume renderingprocessing, and image difference processing as needed. The imagereconstruction process can implement various of the steps of method 100and/or method 100 in addition to various CT image reconstructionmethods. The reconstruction device 514 can use the memory to store,e.g., projection data, reconstructed images, calibration data andparameters, and computer programs.

The reconstruction device 514 can include a CPU (processing circuitry)that can be implemented as discrete logic gates, as an ApplicationSpecific Integrated Circuit (ASIC), a Field Programmable Gate Array(FPGA) or other Complex Programmable Logic Device (CPLD). FPGA or CPLDimplementation may be coded in VHDL, Verilog, or any other hardwaredescription language and the code may be stored in an electronic memorydirectly within the FPGA or CPLD, or as a separate electronic memory.Further, the memory 512 can, be non-volatile, such as ROM, EPROM, EEPROMor FLASH memory. The memory 512 can also be volatile, such as static ordynamic RAM, and a processor, such as a microcontroller ormicroprocessor, can be provided to manage the electronic memory as wellas the interaction between the FPGA or CPLD and the memory.

Alternatively, the CPU in the reconstruction device 514 can execute acomputer program including a set of computer-readable instructions thatperform the functions described herein, the program being stored in anyof the above-described non-transitory electronic memories and/or a harddisk drive, CD, DVD, FLASH drive or any other known storage media.Further, the computer-readable instructions may be provided as a utilityapplication, background daemon, or component of an operating system, orcombination thereof, executing in conjunction with a processor, such asa Xenon processor from Intel of America or an Opteron processor from AMDof America and an operating system, such as Microsoft VISTA, UNIX,Solaris, LINUX, Apple, MAC-OS and other operating systems known to thoseskilled in the art. Further, CPU can be implemented as multipleprocessors cooperatively working in parallel to perform theinstructions.

In one implementation, the reconstructed images can be displayed on adisplay 516. The display 516 can be an LCD display. CRT display, plasmadisplay, OLED, LED or any other display known in the art.

The memory 512 can be a hard disk drive, CD-ROM drive, DVD drive, FLASHdrive, RAM, ROM or any other electronic storage known in the art.

As discussed above, methods 100 and 100′ can also be used with positronemission tomography (PET) data or magnetic resonance imaging (MRI) data.Exemplary implementations of MRI and PET scanners are now provided.

Method 100 (100′) can also be implemented using MRI data acquired usingan MRI scanner such as the non-limiting example of the MRI scanner 700shown in FIG. 7 . MRI is an imaging scan method that magneticallyexcites nuclear spins of a subject placed in a magnetostatic field by aradio frequency (RF) pulse having a Larmor frequency thereof, togenerate an image from magnetic resonance signal data generated with theexcitation.

FIG. 7 illustrates a non-limiting example of an exemplary overview of amagnetic resonance imaging (MRI) system 700 according to one or moreaspects of the disclosed subject matter. The MRI system 700 includes agantry 710 (shown in schematic cross section) and various related systemcomponents interfaced therewith. At least the gantry 710 is typicallylocated in a shielded room. One MRI system geometry depicted in FIG. 7includes a substantially coaxial cylindrical arrangement of the staticfield B₀ magnet 712, a Gx, Gy, and Gz gradient coil set 714 and a largewhole body RF coil (WBC) assembly 716. The physical Gx, Gy, and Gzgradient axes can be controlled in such a way to create G_(RO), G_(PE),and G_(SS) (readout, phase encode, slice-selection) functional axes.Along the horizontal axis of the cylindrical array of elements is animaging volume 718 shown as substantially encompassing the head of apatient 709 supported by a patient table 711. A smaller RF coil 719 isshown as more closely coupled to the head of the patient 709 in imagevolume 718. RF coil 719 can be a surface coil or array or the like, andcan be customized or shaped for particular body parts, such as skulls,arms, shoulders, elbows, wrists, knees, legs, chests, spines, etc. AnMRI system controller 722 interfaces with MRI sequence controller 730,which, in turn controls the Gx, Gy, and Gz gradient coil drivers 732, aswell as the RF transmitter 734 and the transmit/receive switch 736 (ifthe same RF coil is used for both transmission and reception). The MRIsequence controller 730 includes suitable program code structure 738 forimplementing data acquisition sequences including a fast spin echo (FSE)pulse sequence with a time-shifted G_(SS) gradient, for example. The MRIsystem controller 722 also can optionally interface with a printer 728,a keyboard 726, and a display 724.

The various related system components include an RF receiver 740providing input to data processor 742, which is configured to createprocessed image data, which is then sent to display 724. The MRI dataprocessor 742 is also configured for access to previously acquired dataacquisitions of pulse sequences with a time-shifted G_(SS) gradientstored in MRI image memory 746, and to perform various steps of method100 and/or method 100′ stored in code structure 750, as well as MRIimage reconstruction program code structure 744.

Also illustrated in FIG. 7 is a generalized depiction of an MRI systemprogram store (memory) 750 where program code structures (e.g., toperform various steps of method 100 and/or method 100′, for defininggraphical user interfaces and accepting operator inputs to the graphicaluser interface, etc.) are stored in non-transitory computer-readablestorage media accessible to the various data processing components ofthe MRI system. The program store 750 may be segmented and directlyconnected, at least in part, to different elements of the variousrelated system components as needed.

Turning now to an implementation using method 100 (100′) using PET dataacquired using the non-limiting example of PET scanner 800 shown in FIG.8A and 8B. In positron emission tomography (PET) imaging, aradiopharmaceutical agent is introduced into the object to be imaged viainjection, inhalation, or ingestion. After administration of theradiopharmaceutical, the physical and bio-molecular properties of theagent cause it to concentrate at specific locations in the human body.The actual spatial distribution of the agent, the intensity of theregion of accumulation of the agent, and the kinetics of the processfrom administration to its eventual elimination are all factors that mayhave clinical significance. During this process, a positron emitterattached to the radiopharmaceutical agent will emit positrons accordingto the physical properties of the isotope, such as half-life, branchingratio, etc. The radionuclide emits positrons, and when an emittedpositron collides with an electron, an annihilation event occurs,wherein the positron and electron are combined (e.g., an annihilationevent can produce two gamma rays (at 511 keV) traveling at substantially180 degrees apart).

To reconstruct the spatio-temporal distribution of the radia-isotope viatomographic reconstruction principles, each detected event ischaracterized to its energy (i.e., amount of light generated), itslocation, and its timing. By detecting the two gamma rays, and drawing aline between their locations, i.e., the line-of-response (LOR), one candetermine the likely location of the original disintegration to generatea line of possible interaction. By accumulating a large number of suchlines and using a tomographic reconstruction process, distribution ofthe radiopharmaceutical agent can be reconstructed. Additional, using,accurate timing (within a hundred picoseconds) a time-of-flight (TOF)calculation can add more information regarding the likely position ofthe event along the LOR. By collecting a large number of events, animage of an object can be estimated through tomographic reconstruction.

PET imaging systems can use detectors positioned across from one anotherto detect the gamma rays emitting from the object. A ring of detectorscan be used in order to detect gamma rays coming from each angle. Thus,a PET scanner can be substantially cylindrical to be able to capture asmuch radiation as possible, which should be isotropic. A PET scanner canbe composed of several thousand individual crystals (i.e., scintillatorelements), which are arranged in two-dimensional scintillator arraysthat are packaged modules with photodetectors to measure the lightpulses from respective scintillation events. The relative pulse energyensured by the photodetectors is used to identify the position of thescintillation event. The length or depth of the crystal will determinehow likely the gamma ray will be captured. One example of ascintillation crystal is LYSO (or Lu_(1.8)Y_(6.2)SiO₅;Ce or LutetiumOrthosilicate). Other crystals can be used.

Using Anger logic and crystal decoding, the source of each scintillationevent can be identified as originating from a particular scintillator. Ascintillation event will generate light initially radiatingisotopically. The spatial distribution of this light may be modified byinteractions with scintillator surfaces and reflectors before beingdetected by the four nearest photodetectors. From the relative pulseenergy measured by each of these four photodetectors, the position ofthe scintillation event relative to the four photodetectors can bedetermined. The formulas for deriving position information from therelative pulse energies of the photodetectors are referred to as Angerarithmetic. These positions can be further refined by generating alookup table from a floodmap in order to assign each scintillator eventto a scintillator element using the lookup table. This process ofmapping from the x- and y-positions obtained using Anger arithmetic todiscrete scintillator elements is referred to as crystal decoding.

FIG. 8A and 8B show it PET scanner 800 including a number of gamma-raydetectors (GRDs) (e.g., GRD1, GRD2, through GRDN) that are eachconfigured as rectangular detector modules. According to oneimplementation, the detector ring includes 40 GRDs. In anotherimplementation, there are 48 GRDs, and the higher number of GRDs is usedto create a larger bore size for the PET scanner 800.

Each GRD can include a two-dimensional array of individual detectorcrystals, which absorb gamma radiation and emit scintillation photons.The scintillation photons can be detected by a two-dimensional array ofphotomultiplier tubes (PMTs) that are also arranged in the GRD. A lightguide can be disposed between the array of detector crystals and thePMTs. Further, each GRD can include a number of PMTs of various sizes,each of which is arranged to receive scintillation photons from aplurality of detector crystals. Each PMT can produce an analog signalthat indicates when scintillation events occur, and an energy of thegamma ray producing the detection event. Moreover, the photons emittedfrom one detector crystal can be detected by more than one PMT, and,based on the analog signal produced at each PMT, the detector crystalcorresponding to the detection event can be determined using Anger logicand crystal decoding, for example.

FIG. 8B shows a schematic view of a PET scanner system having gamma-ray(gamma-ray) photon counting detectors (GRDs) arranged to detectgamma-rays emitted from an object OBJ. The GRDs can measure the timing,position, and energy corresponding to each gamma-ray detection. In oneimplementation, the gamma-ray detectors are arranged in a ring, as shownin FIGS. 8A and 8B. The detector crystals can be scintillator crystals,which have individual scintillator elements arranged in atwo-dimensional array and the scintillator elements can be any knownscintillating material. The PMTs can be arranged such that light fromeach scintillator element is detected by multiple PMTs to enable Angerarithmetic and crystal decoding of scintillation event.

FIG. 8B shows an example of the arrangement of the PET scanner 800, inwhich the object OBJ to be imaged rests on a table 816 and the GRDmodules GRD1 through GRDN are arranged circumferentially around theobject OBJ and the table 816. The GRDs can be fixedly connected to acircular component 820 that is fixedly connected to the gantry 840. Thegantry 840 houses many parts of the PET imager. The gantry 840 of thePET imager also includes an open aperture through which the object OBJand the table 816 can pass, and gamma-rays emitted in oppositedirections from the object OBJ due to an annihilation event can bedetected by the GRDs and timing and energy information can be used todetermine coincidences for gamma-ray pairs.

In FIG. 8B, circuitry and hardware is also shown for acquiring, storing,processing, and distributing gamma-ray detection data. The circuitry andhardware include: a processor 870, a network controller 874, a memory878, and a data acquisition system (DAS) 876. The PET imager alsoincludes a data channel that routes detection measurement results fromthe GRDs to the DAS 876, a processor 870, a memory 878, and a networkcontroller 874. The data acquisition system 876 can control theacquisition, digitization, and routing of the detection data from thedetectors. In one implementation, the DAS 876 controls the movement ofthe bed 816. The processor 870 performs functions includingreconstructing images from the detection data, pre-reconstructionprocessing of the detection data, and post-reconstruction processing ofthe image data, as discussed herein.

The processor 870 can be configured to perform various steps of method100 and/or 100′ described herein. The processor 870 can include a CPUthat can be implemented as discrete logic gates, as an ApplicationSpecific Integrated Circuit (ASIC), a Field Programmable Gate Array(FPGA) or other Complex Programmable Logic Device (CPLD). An FPGA orCPLD implementation may be coded in VHDL, Verilog, or any other hardwaredescription language and the code may be stored in an electronic memorydirectly within the FPGA or CPLD, or as a separate electronic memory.Further, the memory may be non-volatile, such as ROM, EPROM, EEPROM orFLASH memory. The memory can also be volatile, such as static or dynamicRAM, and a processor, such as a microcontroller or microprocessor, maybe provided to manage the electronic memory as well as the interactionbetween the FPGA, or CPLD and the memory.

Alternatively, the CPU in the processor 870 can execute a computerprogram including a set of computer-readable instructions that performvarious steps of method 100 and/or method 100′, the program being storedin any of the above-described non-transitory electronic memories and/ora hard disk drive, CD, DVD, FLASH drive or any other known storagemedia. Further, the computer-readable instructions may be provided as autility application, background daemon, or component of an operatingsystem, or combination thereof, executing in conjunction with aprocessor, such as a Xenon processor from Intel of America or an Opteronprocessor from AMD of America and an operating system, such as MicrosoftVISTA, UNIX, Solaris, LINUX, Apple, MAC-OS and other operating systemsknown to those skilled in the art. Further, CPU can be implemented asmultiple processors cooperatively working in parallel to perform theinstructions.

The memory 878 can be a hard disk drive, CD-ROM drive, DVD drive, FLASHdrive, RAM, ROM or any other electronic storage known in the art.

The network controller 874, such as an Intel Ethernet PRO networkinterface card from Intel Corporation of America, can interface betweenthe various parts of the PET imager. Additionally, the networkcontroller 874 can also interface with an external network. As can beappreciated, the external network can be a public network, such as theInternet, or a private network such as an LAN or WAN network, or anycombination thereof and can also include PSTN or ISDN sub-networks. Theexternal network can also be wired, such as an Ethernet network, or canbe wireless such as a cellular network including EDGE, 3G and 4Gwireless cellular systems. The wireless network can also be WiFi,Bluetooth, or any other wireless form of communication that is known.

While certain implementations have been described, these implementationshave been presented by way of example only, and are not intended tolimit the teachings of this disclosure. Indeed, the novel methods,apparatuses and systems described herein may be embodied in a variety ofother forms; furthermore, various omissions, substitutions and changesin the form of the, methods, apparatuses and systems described hereinmay be made without departing from the spirit of this disclosure.

1-20. (canceled)
 21. A medical image processing apparatus comprisingprocessing circuitry configured to apply a neural network to atwo-dimensional medical image data to obtain a processed two-dimensionalmedical image data, the two-dimensional medical image data beingacquired by imaging a subject, and the neural network being trainedusing multiple medical image data sets for reducing image noise orartifact, wherein the processing circuitry is configured to inputthree-dimensional volumetric image data comprising the two-dimensionalmedical image data to the neural network to output the processedtwo-dimensional medical image data.
 22. The medical image processingapparatus according to claim 21, wherein the processing circuitry isconfigured to obtain the processed two-dimensional medical image data bycollectively inputting the two-dimensional medical image data and one ormore two-dimensional medical image data sets close to thetwo-dimensional medical image data to the neural network.
 23. Themedical image processing apparatus according to claim 21, wherein thetwo-dimensional medical image data comprises a slice image of thesubject, and the processing circuitry is configured to obtain theprocessed two-dimensional medical image data by inputting the sliceimage and one or more spatially close slice images to the neuralnetwork.
 24. The medical image processing apparatus according to claim21, wherein the two-dimensional medical image data comprises a sliceimage of the subject, and the processing circuitry is configured toobtain the processed two-dimensional medical image data by inputting theslice image and one or more adjacent slice images to the neural network.25. The medical image processing apparatus according to claim 21,wherein the two-dimensional medical image data comprises a slice imageof the subject, and the processing circuitry is configured to obtain theprocessed two-dimensional medical image data by inputting the sliceimage and above and below slice images to the neural network.
 26. Themedical image processing apparatus according to claim 21, wherein thetwo-dimensional medical image data comprises CT image data acquired froma CT scanning on the subject.
 27. The medical image processing apparatusaccording to claim 21, wherein the two-dimensional medical image datacomprises MRI image data.
 28. The medical image processing apparatusaccording to claim 21, wherein the neural network comprises an inputlayer, an output layer, and an intermediate layer between the inputlayer and the output layer, and the neural network is acquired from adeep learning process using the multiple medical image data sets. 29.The medical image processing apparatus according to claim 21, whereinthe two-dimensional medical image data comprises reconstructed imagedata acquired from a reconstruction process based on scan data, the scandata being acquired by scanning the subject using a medical imager. 30.The medical image processing apparatus according to claim 21, whereinthe processing circuitry is configured to output the processedtwo-dimensional medical image data comprising reduced noise or artifactof the two-dimensional medical image data.
 31. A medical imageprocessing method comprising applying a neural network to atwo-dimensional medical image data to obtain a processed two-dimensionalmedical image data, the two-dimensional medical image data beingacquired by imaging a subject, and the neural network being trainedusing multiple medical image data sets for reducing image noise orartifact, wherein the applying the neural network to the two-dimensionalmedical image data comprises inputting three-dimensional volumetricimage data comprising the two-dimensional medical image data to theneural network to output the processed two-dimensional medical imagedata.